Dosimetric verification of radiation therapy is a very important procedure, as successful radiation therapy requires an accurate delivery of dose to a cancerous volume of tissue. Accordingly, when treating patients with radiation, it is desirable to quantify the dose of radiation that will be applied and to verify that such quantified dose will be delivered by the equipment to be utilized. Further, it is believed that a decrease of 10 to 15% in dose delivery will result in a decrease in the chance of cure by a factor of two or three times, while an increase in dose will similarly increase the chance of irreversible damage. Therefore, accurate and specific dose levels are critical to the success rate of patient treatment.
One such method of treating patients is with x-radiation Intensity Modulated Radiation Therapy (IMRT). With IMRT, the radiation is delivered as many tiny, pencil-thin radiation beams (i.e. beamlets), with characterized weights in radiation dose for each beamlet, wherein the beams enter the body from many angles to destroy cancer cells. This accurate delivery of beamlets permits a higher dose of radiation to be delivered to tumors and limits the dose to surrounding healthy tissue, thereby reducing radiation side effects. In this fashion, IMRT can be utilized to safely treat tumors including, for exemplary purposes only, those located near critical organs, such as the eyes and the spinal cord.
To implement treatment, IMRT is computationally planned, based on the computed tomography image of a patient. Computed tomography, or CT, is an x-ray diagnostic procedure utilized to generate a three-dimensional image of a patient, wherein the resulting image is composed of a multitude of cross-sectional views. CT requires data acquisition, image reconstruction and image display. To collect data, x-rays are passed through a patient and are attenuated within the patient, wherein the resulting levels of x-rays are sensed by external detectors to allow the creation of a detailed image of the internal composition of the patient. By moving the x-ray source and taking multiple images, detailed cross-sections can be produced, which then can be utilized to form a three-dimensional image of the patient for accurate selection of the target area to be treated. Once the CT image has been formed and the target area position elucidated and selected, IMRT optimization (in weights and positions of beamlets) is performed to achieve the desired dose distribution in the target and surrounding organs.
Prior to performing IMRT, verification of the radiation levels for the therapy is performed. Typically, a “phantom”, or tissue mimic, is utilized to assist in such verification. First, beams that were optimized for patient treatment are delivered upon a flat phantom, and the consequent dose distributions at some specific depth are calculated for each beam. Second, beams are actually delivered on a phantom that houses a planar dosimeter under a medical linear accelerator, thereby generating signals on the dosimeter. The planar signals are converted into a dose distribution, which is then compared with the dose distribution obtained by calculation. If the difference between the calculations and the actual measurements are within acceptable parameters, the treatment based on the computationally-optimized beam commences on a patient.
As stated above, IMRT requires pretreatment verification based on dose measurement, taken by applying radiation to a phantom that simulates the human organ/tissue. IMRT delivers a radiation dose conforming to the volume of a target only, thus saving exposure of normal and/or critical organs to radiation. In so doing, IMRT generates a non-uniform dose distribution that potentially has a rapid dose gradient. This feature of IMRT puts more demand on multi-dimensional dosimeters, requiring fine spatial resolution and dose integration, in addition to the traditional properties of tissue (or water)-equivalent dose response and real-time data acquisition.
Existing methods of dosimetry include the use of instruments such as ionization chambers (IC), thermoluminescent dosimeters (TLD), and diode detectors. An IC is universally regarded as the standard dosimeter for calibration and dose measurement for radiation therapy, and an IC/water phantom system has been recommended for isodose distribution measurement. However, IC/water phantom systems have some shortcomings in operation. Measurement with an IC provides only selective information with poor resolution. That is, each point-wise datum is limited by its volume and the spacing between measurements. In addition, measurement time using IC/water phantom systems is relatively long. Furthermore, dynamic beam-defining multileaf-collimators and wedges have complicated measurement of doses, taken utilizing IC/water phantom systems.
For dynamic beam dosimetry, a large array IC/water phantom system must be utilized to simultaneously measure doses at various positions in a phantom. However, in addition to the economic disadvantage, simultaneously placing a large number of ICs in a phantom alters the dose distribution being measured and limits spatial resolution (currently-used dosimeters typically have spacing of 7 mm). Such inadequacy similarly applies to TLDs and diode detectors. Therefore, improved dosimeters are needed.
In an effort to overcome the aforementioned deficiencies, and develop a multidimensional dosimeter, prior designs have included use of a TLD plate, plastic scintillators, radiographic (x-ray) film, tissue-equivalent gel, and electronic portal imaging devices (EPID). For instance, a TLD plate utilizing tissue-equivalent elements has been tried. Unfortunately, the diameter of the laser beam reader (i.e., 1.7 mm) limited the spatial resolution of the TLD plate, wherein for isodose measurement of a cobalt-60 radiotherapy beam in a water phantom, the plate showed a relatively high deviation of 14% from measurements utilizing ICs. Such a deviation suggests that utilizing a TLD plate dosimeter requires technically-challenging fabrication of a relatively large plate-type TLD with uniform thickness, in addition to stabilizing material properties, particularly against embrittlement and laser heating.
The potential of utilizing a plastic scintillator for two-dimensional dose measurement of megavoltage beams was first realized through incorporating a digital camera as a light receptor. Other attempts have utilized a plastic scintillator combined with an optical fiber and a photomultiplier tube for brachytherapy dose measurement. Material properties of the scintillator have even been modified to suit dose measurement of low-energy brachytherapy sources. However, no effort utilizing scintillators has been directed to photon beam measurement of IMRT.
Thin radiochromic film has been investigated and utilized as a dosimeter for brachytherapy and less frequently for external radiotherapy that includes stereotactic radiotherapy and IMRT. Such film has relatively good tissue equivalence. In addition to the excellent features inherent in thin film dosimeters, no post-irradiation processing is required for radiochromic film. However, radiochromic film has several undesirable features as a dosimeter: (1) radiochromic film is much more expensive than radiographic film, (2) the optical density of the film after exposure changes with time, which requires an additional calibration of optical density to time, (3) inherent error of uncertainty per single sheet of the film is relatively large (e.g., 1-5% after noise reduction) compared with that of radiographic film (e.g., 0.5% without any processing), and (4) due to the low sensitivity of radiochromic film, a relatively large dose (a few grays (Gy)) is required for a longer exposure time than is required with radiographic film (a few cGy).
Radiographic (x-ray) film has been utilized extensively for IMRT. This is because radiographic film (1) has excellent spatial resolution limited mainly by the grain size in the film emulsion (less than 1 micron) and the aperture of the light beam in a densitometer, (2) has a uniform sensor thickness and response across the film (<0.5% difference on films processed simultaneously), (3) requires relatively short measurement time (less than a few hours including experiment and dose acquisition), (4) is currently inexpensive (e.g., a few dollars for a single sheet of KODAK X-OMAT film—KODAK and X-OMAT are registered trademarks of Eastman Kodak Company), (5) integrates the dose and is thus suited for dynamic beam measurement, and (6) involves a relatively simple measurement setup and procedure, involving a plastic water phantom rather than a liquid water phantom. However, radiographic film suffers from energy-dependent response (and thus inaccuracy) as well as requiring post-irradiation processing (i.e., real-time data acquisition is difficult).
Gel dosimetry using magnetic resonance (MR) imaging technique or optical computed tomography has been studied for three-dimensional work. This technique can provide accurate dose measurements since it employs a practical tissue-equivalent medium that can be molded into a desired shape for use as a dosimeter. However, gel dosimetry is associated with some or all of the following technical difficulties: unavailability of a dedicated MR unit, relatively high expense associated with its operation, difficulties in gel handling, non-uniformity of a magnetic field, variability in the gel manufacturing process, and so on.
Electronic portal imaging devices (EPID) have been investigated for dose verification. Dosimetry utilizing an EPID is normally compared to calculations, using a treatment planning system for verification. However, calculations from a model-based planning system cannot accurately reproduce the exit dose measured in an aSi-based EPID with a scintillation screen containing a Gd-compound. This is due to the over-response of the compound to low-energy photons, unlike the liquid ion chamber-based EPID. In addition, modeling back-scattered photons entering into the aSi-based EPID is not an easy task; although, exit doses have previously been modeled on an aSi-based EPID accurately utilizing a Monte Carlo technique when compared to a measured dose. Furthermore, the calculation suffered from long calculation time (e.g., it took one hour for one IMRT field; thus, for multiple IMRT fields it would require several hours). Modeling all the necessary structural components and back-scattering would be extremely complex, time consuming and difficult. Therefore, exit dosimetry using the aSi-based EPID is still in the developmental stage. Additionally, EPID is in essence an exit dosimeter rather than an in-phantom dosimeter, whereas an in-phantom dosimeter more actually simulates measurement in a patient.
None of the above dosimeters meet the demand necessitated by IMRT. Therefore, it is readily apparent that there is a need for an improved dosimeter in radiation therapy. As will be more fully detailed hereinbelow, it is to the provision of such an apparatus and method that the present invention is directed.